Method and Apparatus for Assisting a Heart

ABSTRACT

An apparatus for a heart of a patient having a cardiac assist device adapted to be implanted into the patient to assist the heart with pumping blood. The apparatus has a sensor adapted to be implanted into the patient. The sensor in communication with the cardiac assist device and the heart which measures native volume of the heart. Alternatively, the sensor monitors the heart based on admittance while the cardiac assist device. Alternatively, the sensor monitors the heart based on impedance.

FIELD OF THE INVENTION

The present invention is related to measuring native volume of a heartof a patient having a cardiac assist device. Volume per time is flow,and typically defined as cardiac output (CO). The term, Native Flow, isnot a standard term, but describes the difference between diastolic andsystolic volumes as contributed by the contraction of the myocardiumitself. The Native Cardiac Output is Native Volume times heart rate. Ina patient without a cardiac assist device. Native CO is equal to totalCO, and native volume is equal to stroke volume. However, in a patientwith a pump, total CO=native CO+pump CO. (As used herein, references tothe “present invention” or “invention” relate to exemplary embodimentsand not necessarily to every embodiment encompassed by the appendedclaims.) More specifically, the present invention is related tomonitoring a heart of a patient having a cardiac assist device withadmittance or impedance.

BACKGROUND OF THE INVENTION

This section is intended to introduce the reader to various aspects ofthe art that may be related to various aspects of the present invention.The following discussion is intended to provide information tofacilitate a better understanding of the present invention. Accordingly,it should be understood that statements in the following discussion areto be read in this light, and not as admissions of prior art.

Cardiomyopathy is a disease of the heart muscle that can lead tocardiogenic shock, a life-threatening condition in which the heart isunable to pump enough blood to support the body's vital organs. In theU.S. alone, cardiomyopathy causes 1.8 million hospitalizations per yearand carries a 30% one-year mortality rate after hospital admission (3).Cardiomyopathy has annual Medicare costs of approximately $20 billion(4) and is the number one cause of hospitalizations and length of stayin patients greater than 65 years old (5). The incidence of cardiogenicshock is increasing, with a >2× increase in the number of cardiomyopathydischarges complicated by cardiogenic shock, from 2004-2014 (6).Cardiogenic shock occurs because the weakened heart suddenly cannot pumpenough blood to the rest of the body to sustain it. In these cases, ashort-term mechanical circulatory support (MCS) device can be placed inthe heart to help maintain high forward blood flow while resting(mechanically unloading) the failing heart. These MCS devices are pumpsthat continuously draw blood from the left ventricle through an inletport and expel the blood into the ascending aorta. The MCS can beinserted via a standard catheterization procedure through the femoralartery, into the ascending aorta, across the aortic valve, and into theleft ventricle (FIG. 1). Once proper placement has been confirmed, thespeed of the pump is set depending upon patient condition.

The ability of MCS devices to maintain peripheral perfusion whilemechanically unloading the heart has the potential to improve mortalityacross at least three large groups of patients (7): 1. high riskpercutaneous coronary intervention (PCI), 2. acute myocardial infarction(MI) with or without cardiogenic shock, and 3. acute decompensated heartfailure.

While the first category of high-risk PCI is an elective surgery and istypically performed over a surgical (short) period, for the last twocategories, MCS devices are implanted for up to 6 days as abridge-to-recovery or bridge-to-decision for a more long-termventricular assist implant. In these patients, indwelling time islonger, and bridging to recovery is not guaranteed.

This is a vast improvement over current practice, whereby a static pumpflow is set for the MCS at implant and does not change until thephysician decides to initiate the judgement-based and arduous deviceremoval weaning process. There are several risks to the patient inprolonging the initiation of this device removal process includinglonger recovery time and sepsis.

BRIEF SUMMARY OF THE INVENTION

The present invention pertains to an apparatus for a heart of a patient.The apparatus comprises a cardiac assist device adapted to be implantedinto the patient to assist the heart with pumping blood. The apparatuscomprises one or more sensors adapted to be implanted into the patient.The sensor(s) in communication with the cardiac assist device and theheart which measures native volume of the heart. The apparatus may beused with a patient during recovery or in high risk, such aspercutaneous coronary, intervention.

The present invention pertains to an apparatus for a heart of a patient.The apparatus comprises a cardiac assist device adapted to be implantedinto the patient to assist the heart with pumping blood. The apparatuscomprises one or more sensors adapted to be implanted into the patient.The sensor(s) in communication with the cardiac assist device and theheart which monitors the heart based on admittance while the cardiacassist device is in operation.

The present invention pertains to an apparatus for a heart of a patient.The apparatus comprises a cardiac assist device adapted to be implantedinto the patient to assist the heart with pumping blood. The apparatuscomprises one or more sensors adapted to be implanted into the patient.The sensor(s) in communication with the cardiac assist device and theheart which monitors the heart based on impedance while the cardiacassist device is in operation.

The present invention pertains to a method for treating a heart of apatient. The method comprises the steps of pumping blood of the patientwith a cardiac assist device implanted into the patient. There is thestep of measuring native volume of the heart with one or more sensorsimplanted into the patient, the sensor(s) in communication with thecardiac assist device and the heart.

The present invention pertains to apparatus for a heart of a patient.The apparatus comprises a cardiac assist device adapted to be implantedinto the patient to assist the heart with pumping blood. The apparatuscomprises one or more sensors adapted to be implanted into the patient.The sensor(s) producing a source signal. The sensor(s) in communicationwith the cardiac assist device and the heart which monitors the heartwith the source signal. The sensor dynamically shifting the sourcesignal to avoid noise from the pump or other sources.

The present invention pertains to a method for assisting a heart of apatient. The method comprises the steps of producing a source signal byone or more sensors implanted in the patient. The sensor(s) incommunication with a cardiac assist device implanted into the patient toassist the heart with pumping blood and the heart. There is the step ofmonitoring the heart by the sensor with the source signal. There is thestep of the sensor dynamically shifting the source signal to avoid noisefrom the pump or other sources.

BRIEF DESCRIPTION OF THE SEVERAL VIEWS OF THE DRAWING

In the accompanying drawings, the preferred embodiment of the inventionand preferred methods of practicing the invention are illustrated inwhich:

FIG. 1 shows the claimed invention in conjunction with the heart.

FIG. 2 is a block diagram of the claimed invention.

FIG. 3 shows the claimed invention.

FIG. 4 shows the cardiac assist device and electrodes in conjunctionwith the heart.

FIG. 5 shows a graph of frequency versus signal strength in regard to asimulated admittance measurement in a water bath with an Impella CP atminimum speed of 23,000 RPM.

FIG. 6a shows an unmodified Impella device.

FIG. 6b shows a modified silver tape Impella device with electrodes.

FIG. 6c shows a modified stainless steel Impella device with fourelectrodes.

FIG. 7 shows a Fourier analysis of motor electromagnetic noise versusadmittance signal for all 9 motor speeds.

FIG. 8 shows in vitro testing with motor at various levels showing lownoise.

FIG. 9 shows the claimed invention.

FIG. 10 shows the claimed invention.

FIG. 11 shows the claimed invention with alternative placement of theelectrodes.

FIG. 12 shows a flexible printed circuit board wrapped around thecatheter.

FIG. 13 shows the embedded system is interfaced to the main computerwith a cable.

DETAILED DESCRIPTION OF THE INVENTION

Referring now to the drawings wherein like reference numerals refer tosimilar or identical parts throughout the several views, and morespecifically to FIGS. 1, 2 and 3 thereof, there is shown an apparatus 10for a heart 12 of a patient. The apparatus 10 comprises a cardiac assistdevice 14 adapted to be implanted into the patient to assist the heart12 with pumping blood. The apparatus 10 comprises one or more sensors 16adapted to be implanted into the patient. The sensor(s) 16 incommunication with the cardiac assist device 14 and the heart 12 whichmeasures native volume and pressure of the heart 12. The apparatus 10may be used with a patient during recovery or in high-risk percutaneouscoronary intervention.

The sensor 16 may include electrodes 18 directly attached to the cardiacassist device 14 that produce signals which are used to measure thenative volume and pressure of the heart 12. The cardiac assist device 14may have a shaft 20 that is adapted to be positioned in the heart 12,and the electrodes 18 are in contact with the shaft 20 that ispositioned in the heart 12 chamber. The electrodes 18 should be in thechamber of interest for the sensor 16 to work properly. If theelectrodes 18 are not in the chamber of interest, the sensor 16 mostlikely will not work.

The sensor 16 may include a computer 22 for data acquisition andanalysis of the signals. The computer 22 is in communication with theelectrodes 18. The computer 22 may provide electrical currents to theelectrodes 18 and may measure corresponding voltages to makeadmittance-based measurements and analyze the admittance-basedmeasurements to make real-time volume and pressure measurements of theheart 12. The sensor 16 may include wiring 24 that is in direct contactwith the electrodes 18 and which extends to the computer 22 over whichthe electrical currents pass creating the corresponding voltages. Theremay be a pressure sensor 51 adapted to be implanted into the patient.The pressure sensor 51 in communication with the cardiac assist device14 and the heart and the computer 22 which monitors the left ventricularpressure while the cardiac assist device 14 is in operation. Inaddition, the pressure sensor with the computer may plot native pressurevolume loops while the cardiac assist device is in operation, forinstance, to measure the work done by the heart and its efficiency. Aconsiderable amount of information on cardiac performance can bedetermined from the pressure vs. volume plot.

The cardiac assist device 14 may include a motor 26 and an impeller 28disposed in the shaft 20 which is driven by the motor 26 to assist theheart 12 with pumping blood. The cardiac assist device 14 may have amarker 30 to guide proper placement of the cardiac assist device 14 inthe heart 12. The electrodes 18 may be disposed on the shaft 20 of thedevice. Alternatively, the apparatus 10 may include a catheter having ashaft 20 disposed alongside the shaft 20 of the device and theelectrodes 18 may be disposed alongside the shaft 20.

The cardiac assist device 14 may be a temporary mechanical circulatorysupport (MCS) device which is a catheter-mounted blood pump that drawsblood from a left ventricle of the heart 12 through an inlet port 34 ofthe MCS 32 and expels blood into an ascending aorta 36 of the heart 12,thereby reducing some of the mechanical load on the heart 12 andpromoting recovery. The pump may draw the blood intermittently in apulsatile manner that mimics the natural pulsatile movement of the heart12 or the pump may draw the blood from the left ventricle continuously.If the pump draws the blood intermittently, the timing of theintermittent action and the pulsing of the pump blood may be coordinatedwith the pumping motion of the heart 12. The native cardiac output (CO)and pressure measurements by the sensor 16 may provide a feedback signalto the MCS 32 to modulate flow/volume by the MCS 32 during treatment.

In contrast to having utility for weaning and treatment in a temporaryMCS device that is catheter-mounted and connected to a mains poweredexternal controller, a permanent implant battery-powered MCS device(often used for bridge-to-transplant rather than bridge-to-recovery),would benefit from a feedback signal to control pump speed in differentscenarios. These include battery management, demand feedback, andlong-term dysfunction diagnosis, for example, battery management isimportant in implanted devices, whether rechargeable or not, becausepatient's quality of life is typically tied to how unobtrusive thedevice is to the activities of everyday living. In general, the fasterthe pump flow, the more power is necessary to produce that flow. Totalflow is the important metric to sustain in patients with an implantableMCS device, so having a high pump flow when the native output is highunnecessarily wastes battery life, directly leading to lower quality oflife.

Demand feedback is typically unimportant in temporary MCS scenarios,because they are typically indicated for rescue, or high-risk situations(like high-risk PCI). In these emergent situations, the pump flow istypically set for as high as the patient will tolerate, because largertotal flow for short periods is protective, while total flow that isslightly low can be extremely detrimental due to concomitant conditions.In long-term implant scenarios, the patient will exhibit a larger rangeof possible flow demand, for example, when exercising, for which ahigher-than-normal pump flow is necessary. Other devices like demandpacers (https://www.biotronik.comlen-us/products/services/cls) measureand respond to this demand by using end systolic volume (the minimumvolume sensed) as a surrogate for contractility, which is a marker forblood flow demand. This technique of demand measurement has theadvantage of not being tied directly to activity, which could easily bemeasured using accelerometers. One example of non-activity relateddemand is changing pump flow to meet the demands on the body created byintense emotion (anger typically requires higher blood flows, and alsoincreases contractility on a beat-to-beat basis, for example). Demandpacers can increase cardiac output by increasing heart rate, but a longterm MCS device could directly modulate higher blood flow (strokevolume) instead.

Long-term dysfunction due to remodeling changes in the heart (usuallymarked by an increase in end-diastolic volume over time) are too long ofa time scale to matter in temporary MCS devices. However, in completelyimplanted MCS devices, the native heart measurement can be used todetect worsening (or improving) heart failure status purely as adiagnostic.

The present invention pertains to an apparatus 10 for a heart 12 of apatient. The apparatus 10 comprises a cardiac assist device 14 adaptedto be implanted into the patient to assist the heart 12 with pumpingblood. The apparatus 10 comprises one or more sensors 16 adapted to beimplanted into the patient. The sensor(s) 16 in communication with thecardiac assist device 14 and the heart 12 which monitors the heart 12based on admittance while the cardiac assist device 14 is in operation.

The present invention pertains to an apparatus 10 for a heart 12 of apatient. The apparatus 10 comprises a cardiac assist device 14 adaptedto be implanted into the patient to assist the heart 12 with pumpingblood. The apparatus 10 comprises one or more sensors 16 adapted to beimplanted into the patient. The sensor(s) 16 in communication with thecardiac assist device 14 and the heart 12 which monitors the heart 12based on impedance while the cardiac assist device 14 is in operation.

The present invention pertains to a method for treating a heart 12 of apatient. The method comprises the steps of pumping blood of the patientwith a cardiac assist device 14 implanted into the patient. There is thestep of measuring native volume of the heart with one or more sensors 16implanted into the patient. The sensor(s) 16 in communication with thecardiac assist device 14 and the heart 12. The method can also, oralternatively, use the sensor 16 and the cardiac assist device in thevarious embodiments described herein.

In the operation of the invention, a temporary MCS 32 device is acatheter-mounted blood pump, typically placed for less than 7 days inpatients with cardiomyopathy and cardiogenic shock, that continuouslydraws blood from the left ventricle through an inlet port 34 of the MCS32 and expels the blood into the ascending aorta 36, thereby reducingsome of the mechanical load on the heart 12 and promoting recovery(hemodynamic support). An MCS 32 is unlike other more permanently placedpumping devices, because it is placed in the patient using a standardcatheterization procedure (without piercing the heart 12). Thistechnique is preferable for use as a bridge-to-recovery, because it canbe easily removed. Once proper placement has been confirmed, the speedof the pump is set depending upon patient condition on a case-by-casebasis, and treatment is deemed complete once the heart 12 has recoveredand hemodynamics have returned to normal.

Initiating the device removal weaning process is based upon numerousfactors, most importantly Cardiac Power Output (CPO=CO*Mean AtrialPressure). During continuous use of an MCS 32 device, the total CO tothe body comes from two components: total CO=“native CO”+“device CO”.“Native CO” is the amount of blood ejected from the ventricle by therecovering heart 12, and “device CO” is the amount of blood pumped bythe MCS 32 device. Over time, the native CO from the heart 12 naturallyrises, as it recovers slowly from decreased load, signaling improvement.Removal of the MCS 32 device requires “weaning” the patient from pumpsupport by reducing pump speeds prior to removal. Optimal MCS 32 usewould require reducing pump speed over an extended time as the native COreaches near-normal levels, but in practice native CO is never measuredduring recovery, because it is clinically impractical to continuouslymeasure CO using echocardiograms during the entire recovery period(could be days).

Ideally, an MCS 32 device would continuously measure real-time nativeCO, and then automatically adjust the pump speed to modify the deviceCO, while maintaining an overall total CO. This method would naturallywean the patient as he/she recovers as the native CO rises. Thephysician could monitor the native CO measurement and remove the MCS 32device when appropriate. In Phase I, BSM proved that accuratemeasurements of real-time native CO are possible using anadmittance-based Impella prototype (“Impella-CO”) while operating withinelectromechanical pump noise. This was designed and demonstrated both onthe bench, and in an animal model showing excellent agreement withmultiple CO standards.

Admittance measurement is blocked by electrical insulators like the pumpbody (made of plastics). It can only tell what the volume of the outsideblood pool is, because the measurement relies on the flow of electricitybeing “admitted.” Total CO is very easy to measure, and can be doneusing a number of different methods (Fick, Thermodilution, etc.). NativeCO of the left ventricle can only be determined by imaging methods(directly visualizing the size of the pumping chamber, the LV), andAdmittance. As shown in FIGS. 1 and 4, the outlet port 38 of the deviceis in the ascending aorta 36, outside of and downstream from the chamberof the heart 12. In this way blood from the outlet port 38 mixes withthe native output of blood from the heart 12, downstream of where bloodpumped solely by the heart 12 leaves the heart 12, so only the Native COis measured by admittance. The blood inside the device is not measuredbecause the pump body is made of plastic and shields the blood in thepump from being measured.

The pump typically does not directly measure the amount of flow goingthrough it, and can only estimate its own flow (not usually measure itdirectly) by using the power delivered to/consumed by the pump as asurrogate for how hard the pump is working to pump blood (assuming theinlet and outlet remain unblocked).

Native Cardiac Output is the amount of forward blood flow that is due tothe pumping action of the heart 12. To clarify, when there is no pumppresent, Cardiac Output is equal to Native Cardiac Output. When there isa pump present, the pump flow (the flow of blood through the pump) inL/min is the pump output. Then, the total Cardiac Output is equal to thepump flow plus the Native Cardiac Output, Total CO=Pump CO+Native CO.

Cardiomyopathy is a disease of the heart 12 muscle that can lead tocardiogenic shock, a life-threatening condition in which the heart 12 isunable to pump enough blood to support the body's vital organs. In theU.S. alone, cardiomyopathy causes 1.8 million hospitalizations per yearand carries a 30% one-year mortality rate after hospital admission (3).Cardiomyopathy has annual Medicare costs of approximately $20 billion(4) and is the number one cause of hospitalizations and length of stayin patients greater than 65 years old (5). The incidence of cardiogenicshock is increasing, with a >2× increase in the number of cardiomyopathydischarges complicated by cardiogenic shock, from 2004-2014 (6).Cardiogenic shock occurs because the weakened heart 12 suddenly cannotpump enough blood to the rest of the body to sustain it. In these cases,a short-term mechanical circulatory support (MCS) device can be placedin the heart 12 to help maintain high forward blood flow while resting(mechanically unloading) the failing heart 12. These MCS 32 devices arepumps that continuously draw blood from the left ventricle through aninlet port 34 and expel the blood into the ascending aorta 36. The MCS32 can be inserted via a standard catheterization procedure through thefemoral artery, into the ascending aorta 36, across the aortic valve,and into the left ventricle (FIG. 1). Once proper placement has beenconfirmed, the speed of the pump is set depending upon patientcondition.

The ability of MCS 32 devices to maintain peripheral perfusion whilemechanically unloading the heart 12 has the potential to improvemortality across at least three large groups of patients (7): 1. highrisk percutaneous coronary intervention (PCI), 2. acute myocardialinfarction (MI) with or without cardiogenic shock, and 3. acutedecompensated heart failure.

While the first category of high-risk PCI is an elective surgery and istypically performed over a surgical (short) period, for the last twocategories, MCS 32 devices are implanted for up to 6 days as abridge-to-recovery or bridge-to-decision for a more long-termventricular assist implant. In these patients, indwelling time islonger, and bridging to recovery is not guaranteed.

Preliminary experiments were conducted on test MCS 32 devices providedby Abiomed to determine the initial noise spectrum (FIG. 5). The noisespectrum of the motor 26 vs. the “signal” (the 20 kHz waveform fromCardioVol) was measured over all available pump speeds (0 to 46,000rpm). It was determined that 1) pump noise is essentially the sameregardless of pump speed, but has large spikes at 50, 100, and 150 kHz(this is good because CardioVol can operate at any frequency in the1-100 kHz range, and is currently configured for 20 kHz) and 2) noisespectrum is dependent on the electrode geometry used (also good, becauseunlike traditional conductance, admittance utilizes only four electrodes18 at a time, allowing flexibility in design that allows integrationwith existing devices like the MCS Impella).

FIG. 6 shows a black band located between the wiring 24 exit and themost proximal electrode that is radiopaque and used clinically to alignwith the aortic valve to guide proper placement. This design ensuresthat the surgeon will place all four electrodes 18 in the leftventricular heart 12 chamber.

Motor 26 noise quantification: Fourier frequency analysis of the motor26 noise signal during pump operation was used to determine the optimalfrequencies to use for the cardiac volume measurements. Volumecalculations using admittance-based techniques allows for real-timemeasurement of both blood and muscle contributions by taking advantageof the differing electrical properties in the frequency range ofinterest. Within the frequency band of 1 kHz-100 kHz blood is purelyresistive, but myocardium is both resistive and capacitive. Admittancemeasurements are performed in the complex Fourier plane, which allowsmeasurement and separation of both the capacitive and resistiveproperties of the volume, allowing a pure measurement of blood volume tobe extracted from the total signal. Because the measurement can be madeat any frequency in the span, it is possible to pick a frequency that isuncorrupted by a known noise source, like that of the Abiomed Impellamotor. As shown in 7, the excitation current signal used to make theCardiac Output measurement (FIG. 7, Signal) is several orders ofmagnitude larger than the noise floor (FIG. 7, Noise floor Target). Notethat the pump noise can be clearly seen at 50 khz, and at 100 kHz at all9 pump speeds, and does not overlap with the measurement. It is clearfrom task 1 that any frequency at least 5 kHz away from DC, and not aninteger multiple of 50 kHz could be utilized for the admittance-basedvolume measurement. 20 kHz is chosen for the prototype design tomaximize signal to noise ratio while minimizing power and reducing thecomplexity of a higher frequency design. FIG. 7: In vitro testing withPump at various levels showing low noise. Dynamic testing simulating a“heart beat” by squeezing the sides of a bottle while the pump is at P3,and P6. No filtering has taken place, and the raw admittance (volume)signal is unaffected by motor 26 noise.

Task 2: Modify the operating frequency and spatial electrodeconfiguration of an admittance unit for maximum signal to noise ratiousing information from task 1.

Admittance Unit Design: An instrument to measure admittance derivedblood volume was designed to generate a constant amplitude current at anoperating frequency of 20 kHz injected into the outer two electrodes 18prototyped onto the Impella pump, and used to measure the resultingvoltage from the inner two electrodes 18 as describe in task 1. Becausethe current is constant, the blood volume is proportional to themeasured voltage from the inner two electrodes 18.

In vitro testing: A bottle of saline with conductivity designed to mimicthe conductivity of blood (s=8 mS/cm) was used to determine if theadmittance measurement is sensitive to changing volumes in the presenceof electrical pump noise, and because the bottle has flexible sides, itis possible to modulate the volume sensed near the Impella pump byrepeatedly squeezing the bottle to simulate a beating heart 12. In thisway, a time-varying signal is displayed on the GUI and recorded toexamine signal quality (FIG. 8). The admittance instrument is shown tobe highly responsive to changing volume in the bottle even in thepresence of the pump operating at standard flow levels P3 and P6 (thereare 9 pump speeds total, in addition to “off”, so P3 and P6 representmiddle values of speed). Note that when there is no simulated pumping,the motor 26 turning on does not change the value measured by theprototype impella device, illustrating that the native CO is unaffectedby motor 26 noise. The level of change caused by the pump in astationary setup is less than 1 mQ (milli-Ohm).

Signal to Noise Ratio (SNR) measurement: in vitro, SNR measurements arestraightforward, and can be calculated using a discrete Fouriertransform (DFT) on an oscilloscope. and an ideal signal which is derivedfrom a continuous sine wave. However, in vivo, this measurement is notpractical, given the added power required to maintain the continuoussine wave. Instead, a pulsed current excitation waveform is used to makeextremely fast impedance measurements at 20 kHz. and calculate aFast-Fourier Transform (FFT) in real time. This requires a modified SNRmeasurement approach for accuracy. In general, ReZ and ImZ are used tocalculate the volume signal. Therefore, Signal (S) power can becalculated. The noise (T) can be calculated from all of the other binsof frequency in the FFT. In this way, an SNRa which varies from 0 to1000 (1000 being a “perfect” no noise sine wave, and 0 being a lack ofany discernable signal) can be calculated. Values above 900 areconsidered extremely low noise, and indicate a noise-free measurement inour in vivo measurements. The governing equations are shown below:

S=ReZ ² +ImZ ²(V ²) at 20 kHz

T=Σ _(non 20 kHz bins) Sig−Avg

SNR_(a)=1000*S/T

Task 3: Measure native CO using admittance vs. standards in an animalmodel while simultaneously operating an MCS 32 device (Impella heartpump).

Goals: The goal for this preclinical evaluation was to measure CO usingthe admittance instrument connected to a prototype Abiomed ImpellaMechanical Circulatory Support device and compare those measurements tostandard clinical CO measurement methods. There is currently nouniversally accepted “gold standard” for CO measurement, and while everymethod has its weaknesses, our goal was to show good agreement with atleast two of the current standards. Good agreement was defined aswithin-subject coefficient of variation (wCV)<20%, which is theagreement level that two expert echo readers show when reading the sameset of echoes. 0% would indicate perfect agreement wCV is mathematicallydefined as the standard deviation divided by the mean.

Protocol: N=3 pigs are instrumented with: 1) the modified Impella pumpwith CO measurement capability connected to the admittance unit, and 2)independently measured CO using a Swan Ganz catheter in the pulmonaryartery capable of measuring 2A) CO via thermodilution continuous cardiacoutput, or CCO (Edwards Scientific), 2B) CO via bolus instantaneouscardiac output, or ICO (Edwards Scientific) and 2C) CO via myocardialoxygenation (Fick method). CO using 3) echocardiography with the aorticvelocity time integral method, or aortic VTI at the level of the aortaare also measured. Measurements of CO were taken before the Impella pumpwas implanted, after the pump was implanted in the “off” state, and ateach of the nine pump speeds (P1, P2, P3, P4, P5, P6, P7, P8, and P9).The protocol was discontinued if there was hemodynamic instabilitypresent, indicated by a “suction” alarm from the Impella. In allinstances, “native heart” CO was calculated by taking the total COmeasurement and subtracting the pump flow. The only exceptions were echomeasurements (echo cannot measure the pump flow at the level of theaorta because the pump itself is opaque to sound waves), and admittance(admittance measures only the electrical properties of blood andmyocardium outside the Impella because the catheter body is anelectrical insulator opaque to electrical waves).

Results: In all three pigs, agreement with at least one of the standards(wCV≤16%) and in one of the pigs (Animal 2) were able to be shown,wCV=9% for all 5 standards simultaneously were able to be shown, for allof the Impella pump speeds. Further, while there was no CO changeplanned in the protocol as a result of changing pump speed, the naturaloffloading of the heart 12 created by the Impella device shows a lowernative CO for each progressively higher pump setting. This realphenomenon indicates not only that the Impella device offloads a failingheart 12 while maintaining forward CO and mean arterial pressure, butthat it is possible to make this measurement in real-time using theprototype instrument developed.

Admittance vs. Conductance—There are three novel concepts that must beutilized to solve the problems preventing an Admittance measurement insitu with a generic noise source like a pump, 1) Traditional conductancecatheters (including “dual frequency” devices) cannot determine accuratevolumes because they all subtract a single value for parallelconductance. Using admittance-based technology solves these issues usinga measurement of the capacitive nature of the myocardium, but bothmeasurements are sensitive to noise from the motor 26, so frequency ofoperation must be redesigned. 2) The wires and electronics of anadmittance-based catheter must be extremely small so as not tosubstantially increase the outer diameter of the MCS 32 device, which is12 French in the case of the Abiomed Impella CP, (Danvers, Mass.) and 3)Electrode placement must be designed from scratch around the functionalelements of the Impella device, which affects the field geometry.

General Considerations—In general, conductance measurements of volumeare made by considering the chamber of interest and creating a customcatheter (a conductance catheter) that spans the area to make the mostlinear measurement of impedance possible. A measurement of admittancecan be adapted for use on an arbitrary catheter, which is advantageousbecause many instruments are implanted either temporarily (like apercutaneous heart 12 pump or the cardiac assist device 14, see FIG. 1)or more permanently (like a cardiac implantable electronic device).These instruments often already have electrodes 18, microcontrollers,and signal processing that work in concert to achieve a goal like pacingthe heart 12, or supporting a weak heart 12 by increasing forward flow.The adaptation of an admittance measurement to these existing resourcescan have diagnostic advantages at low additional cost, by carefullyconsidering design tradeoffs to achieve a goal of low barrier to marketentry, or reduction of contaminating noise. In this patent, it isdiscussed how to flexibly achieve both goals to combine a hemodynamicmeasurement with an existing heart 12 pump by taking advantage of twomain ideas: 1) Working with non-ideal electrode locations, geometries,and inter-electrode spacing to linearize a measurement made in anon-linear region, and 2) using a frequency of measurement that avoidsthe noise of nearby. known interference.

Electrode Material—Electrodes 18 should be made of a biocompatiblematerial, that is low resistance, like platinum or gold. The reasoninghere is that the electrode itself should contribute as little aspossible to the final measurement, making calibration less complicated.

Electrode Number—It is widely known that the electrode-electrolyteinterface impedance that arises from putting a metal electrode in thepresence of an electrolyte can change the measured total impedance. Toprotect against this, typically four electrodes (tetrapolar) 18 arenecessary at a minimum, to reduce polarization effects. If numbered fromthe distal to the proximal end of a catheter as 1 (distal tip), 2, 3, 4(most proximal electrode), typically electrodes 1 and 4 provide afunction of current generation by injecting current, and 2 and 3 areused to measure the voltage arising from the current flowing through theelectrolyte. Tetrapolar electrodes provide an advantage, because thedesign of a good constant current source does not change with electroderesistance change, ensuring that the admittance (current over voltage)does not change with fouling of electrodes 1 or 4.

Fouling is a change in resistivity that occurs due to a local change inthe resistance of the electrode. Typically, this is caused by scartissue formation, oxidation, or other change in theelectrode/electrolyte impedance on the voltage electrode, either bychanging the effective geometry, or the affecting the resistivity. Thistetrapolar technique is used often in resistance/impedance measurementsin non-medical fields.

If fewer electrodes 18 are desired or necessary (for example, if onlythree electrodes 18 are present, and used for another purpose, as whenmaking an admittance measurement using a right ventricular implanteddefibrillator lead, with a large shocking coil, ring, and tip conductor[1], [2]), then any electrode connected to both a current generation andvoltage measurement should be the largest possible. This ensures that ifelectrode fouling occurs, it is likely that the change in totalimpedance will not affect the large percentage of the area. For example,R=ρ*UA, and if the A is very large, then a change in resistivity of asmall area Δp/ΔA will not affect the R by very much, even if Δrho islarge.

Electrode Spacing—If the outer two electrodes (1 and 4) are currentgenerating, and the inner two electrodes (2 and 3) are voltagemeasuring, a naïve approach would be to equally space them. However, thesensitivity of a constant-current volume measurement is largest when thelocation of the current and voltage electrodes 18 is close together(that is, 1 and 2 are close together, 3 and 4 are close together), andfar apart from the other pair (2 and 3 are far apart).

Additionally, any admittance change between current and voltage pairs(between 1 and 2, or between 3 and 4) will inversely affect the totaladmittance, a concept called “negative sensitivity”.

As an example: consider a complex impedance measurement. If a highresistance, zero capacitance object enters between electrodes 2 and 3,then the real impedance measured will increase (because of the increasedresistance to current flow, measured as a larger voltage on 2 and 3). Ifthe same object enters between 1 and 2 instead, the real impedance willdecrease (because the same voltage drop occurs between 2 and 3). This isthe concept known as “negative sensitivity”, and was first discussed byLarson et al [3]. In general, electrodes 1 and 2 (and also electrodes 3and 4) should therefore be placed as close together as possible, givenany limitations that must be worked around considering the geometry ofthe catheter shaft 20.

Ideal Electrode Geometry—The ideal physically realizable electrode foran admittance measurement would be a sphere, because it would allow forcurrent output and voltage input from all directions simultaneously.However, to be disposed on a shaft 20 like a catheter, the closest thatcan be achieved is a ring. Typically, single electrodes are ring shaped,and about as long as their diameter. Using a ring shape makes themeasurement independent of the rotational angle of the catheter. Thisallows the connection to the ring to be made within the catheter body(which is usually non-conducting) and keep the wiring 24 harness fromaffecting the impedance measurement. Additionally, it is not generallydesirable to minimize the area of the electrode, because the impedanceof the electrode itself should remain low to avoid affecting themeasurement. This can be done by using a material with extremely lowresistivity (usually at considerable cost if it remains biocompatible)or it can be done more easily by using a larger surface area. It hasbeen found that a length about the size of the diameter is sufficient.In general, the ring size should be thin (the length L should be shorterthan the diameter, d). Additionally, the cross-sectional area π*d*Lshould be large enough that the resistance when placed in normal salineis <10 Ohms. The specific dimensions relate to the material propertiesof the electrode, and the geometry used (thin electrodes with L<<d willhave high resistance. Large d electrodes require a large catheter body,but will have a lower resistance). For the electrodes 18, the diametershould be the same as the catheter tubing diameter. The electrodediameter has a range of 2-5 mm and the electrode width to be 1-3 mm.

Wiring 24 Considerations—Connection of the electrodes 18 to the currentsource and voltage measurement is non-trivial, and in general, long,closely spaced wiring 24 creates a distributed capacitance that must beaccounted for when calibrating. Wiring 24 must also be attached to theelectrodes 18 in some way, ideally from the inside of the electrodering. The catheter body is usually made of a non-conducting material,and if there is no necessary lumen, the wire routing can be madeinternal to the catheter, ensuring that the electrode surface area isnot affected by the wire insulation. However, it is acceptable to routethe wires outside of the catheter body, in particular when the lumenwill need to be used, as in Embodiment 1 below. In these instances, thewiring 24 should be as thin as possible, bonded to the electrode in away that uses minimal external surface area (so as not to affect themeasurement). The wires themselves should be placed farther apart tominimize their effects on capacitance, and should not be coiled so as tominimize inductance effects.

Another potential advantage of having thin wires is that often timescatheters only have lumens because they need to remain patent for theflow of a liquid. Thin wires can more easily avoid impeding this flow ofliquid either physically, or avoid being used as a scaffold forclotting. Additionally, thinner wires are more easily capable of beingbuilt into the lumen walls to completely avoid these effects.

The wires leading to each of the four electrodes 18 create a capacitancebetween them. In general, it is possible to calibrate out if the wiresstay the same distance from one another over the heartbeat. This isbecause changing distance between the wires over the heartbeat could bemisrepresented as change in capacitance measured in myocardium.

Wire attachment is accomplished either by directly soldering the wire tothe electrodes 18 themselves (typically from the inside of the lumenthrough a bored hole) or by using conductive epoxy solder paste throughthe same hole. It is possible to attach electrodes 18 to wires on theoutside, but sensitivity to changes in the area of the solder joint maybe changed by the electrical properties of the wire and attachmentpaste/solder. Electrical properties of the wire and attachmentpaste/solder are different depending on geometry, material, and materialinterface type, and these can all be either modeled using finiteelements, or empirically measured in vitro using prepared saline thathas the conductivity of blood.

The wires used to connect to the electrodes 18 will be 40 to 48 AWG tominimize any increase to the overall outer diameter of the pump. Thewires should be insulated. Embedding the wires in the walls fixes theirposition with respect to one another, providing a constant resistivecontribution. This will help with two factors: A) ease of calibration bykeeping the effects of the wires constant, and B) it would allow keepingthe outer diameter of the pump itself the same (which is an advantagesurgically, because it means you do not need to increase the size of thedelivery sheath or hole in the vessel).

Design for noise-band immunity—Admittance measurement in many differentpotential configurations has been discussed thoroughly in the past. Whenmaking these types of measurements in the presence of noise, analog ordigital filtering can be employed to some success. However, when thenoise band of interference is small, and does not span the entireoptimal range of admittance measurement (10 kHz-100 kHz), [4], a changein current stimulation frequency and an aggressive bandpass filter onthe voltage input can be used to avoid the noise entirely.

There are a few advantages to characterizing the aggressiveness of thenoise in the Fourier domain. The closer a noisy band is to the frequencyof the current source, the harder it is to use a filter, either analogor digital, to separate the signal from the noise. Using high Q analogfilters can be expensive in parts, and high Q digital filters can beexpensive computationally. Delays due to digital filtering at high Q canalso pose problems. Changing the stimulation frequency of the currentgeneration circuit only requires changes in the oscillator that is beingused, allowing for cheaper and faster components to provide the signalprocessing necessary to make a complete measurement.

Improving SNR with dynamic Sin DAC frequency—It is possible to choose asignal generation technique that will allow for small changes in thegenerated current frequency using only software, provided the signalgeneration is driven by a microcontroller clock or other oscillator thatis sufficiently higher than the current generation frequency. In fulldesigns such as those that use high Q analog filters on the voltagesampling section, only small changes in the current frequency generationare useful (less than 5% difference from the initial frequency). becausethe SNR of the returning voltage will suffer from degradation by thefilter. Typically, the waveform is represented in microcontroller memoryas an N-point waveform signal (a sine wave) reproduced point-by-point ona Digital to Analog Converter (a Sin DAC) that has an output rate at ornear the clock frequency. This limits the upper bounds of the Sin DACfrequency output per the following equation:

Sin DAC frequency (Hz)=Clock Speed (Hz)/Length of Waveform N×divider M

Having a higher N value or divider M increases the fidelity andstability of the current outputs at the cost of max Sin DAC frequency,while having a lower clock speed directly reduces the maximum Sin DACfrequency. Generally, there are only a few (10+) discrete speeds atwhich the Sin DAC can operate in this design, but a few hundred Hzdifference may reduce the noise faster than the signal in such a way asto improve the SNR substantially, leading to improvement.

In one embodiment, a change of about 800 Hz in the Sin DAC frequencyover 6 steps was created, by setting the rate of the Sin DAC using theADC interrupt frequency. Given an 80 MHz clock rate, and an N=32 pointsine wave with an M=125 divider, sufficiently clean signals are producedat a Sin DAC frequency of 20 kHz. Changing the divider M changes boththe ADC sampling rate and the output of the N=32 point sine wave,synchronizing both the Sin DAC current and the sampling of the returningvoltage. These outputs would be well within the roll-off of a band-passfilter used to precondition the returning voltage signal, allowing apotential increase in Signal to Noise Ratio.

Dynamic noise-filtering with dynamic Sin DAC frequency—One way to removethe limitations of only a few discrete speeds of Sin DAC is to changethe filtering technique on the voltage inputs by using a mixed-signaldesign switched-capacitor filter network with a high Q (ideally 4-8,given a chip like the Maxim 7491).

Improving SNR while sacrificing signal bandwidth—One way to improve SNRis to collect and process a buffer of data with a longer time interval.For example, if the buffer length is 2 ms, then volume versus time canbe measured at 500 Hz. However, if the buffer length is increased to 20ms, volume versus time is measured at 50 Hz, but SNR is greatlyimproved.

Left ventricular pressure—To improve clinical usefulness of theinvention it is possible to include a pressure sensor 51 within the leftventricle. The challenge is not necessarily the size of the transducer,but the cost of routing three additional wires along the cardiac assistdevice 14. One solution to the device 14 diameter is to embed anembedded computer 53 along with the pressure sensor 51 distal to theinlet port 34 of the cardiac assist device 14. FIG. 12 shows a flexibleprinted circuit board 55 wrapped around the shaft 20. The printedcircuit board 55 has the embedded compute 53 and the pressure sensor 51.FIG. 13 shows the embedded computer 53 and pressure sensor 51 isinterfaced to the main computer 22 with a cable 57 having 3 or morewires. In this configuration only 3 wires are necessary to implementboth admittance-derived volume and left ventricular pressure. One wireis voltage supply, one wire is ground, and the third wire is serialoutput. The serial output encodes the raw measurements, which will thenbe processed in the usual way on the main computer 22. In thisconfiguration the admittance circuitry (sin DAC, and ADC) is included onthe flexible PCB wrapped around the tip of the pump.

Clinical Use Cases: General—The ability of MCS 32 devices to maintainperipheral perfusion while mechanically unloading the heart 12 has thepotential to improve mortality across at least three large groups ofpatients (7): 1. high risk percutaneous coronary intervention (PCI), 2.acute myocardial infarction (MI) with or without cardiogenic shock, and3. acute decompensated heart 12 failure.

While the first category of high-risk PCI is an elective surgery and istypically performed over a surgical (short) period, for the last twocategories, MCS 32 devices are implanted for up to 6 days as abridge-to-recovery or bridge-to-decision for a more long-termventricular assist implant. In these patients, indwelling time islonger, and bridging to recovery is not guaranteed.

Goals of hemodynamic support—The overarching goal of hemodynamic supportis to maximize function by increasing both Cardiac Output (CO, the MCS32 control point), and Mean Arterial Pressure (MAP, optimized bymedication). The product is the Cardiac Power Output (CPO=total CO*MAP)and low CPO is now well accepted as the single most important correlateof mortality in cardiogenic shock. Studies have also shown that the goalof keeping the CPO≥˜0.6 W with MCS 32 devices is predictive ofdecreasing heart failure at 30 days. While the measurement of pressureis currently integrated into many MCS 32 devices (e.g., the Impelladevice has an integrated pressure sensor 16, used for device placement),the only control point for an MCS 32 device is the motor 26 speed, whichincreases CO with increasing motor 26 speed (flow). In practice, thistranslates to a clinical support goal of setting the MCS 32 device flowas high as possible without the causing suction due to high pump speeds(where the motor 26 is spinning, but no blood is moving through the MCS32). The positive impact of an MCS 32 device is highly dependent on theflow rate of the pump, which can be adjusted in the majority of pumps,but the initial flow rate is rarely changed during recovery.

The key indicator of patient stability is Total Cardiac Output (totalCO), which is the summation of the pump CO (the amount of blood inLiters ejected by the pump per min), and the native CO (The amount ofblood in Liters ejected by the heart per min). Total CO=pump CO+nativeCO. Native CO is dependent on patient health, while pump CO iscontrolled by the operator of the MCS 32 device, and estimated using thecurrent draw on the motor 26, or measured RPM, or in some cases, by aflow measurement integral to the pump that is either EM, acoustic, orpressure differential based.

Clinical Use Case 1: Native CO to inform start of weaning from MCS32—MCS 32 device implantation is not without risk, and longer indwellingtimes can lead to infectious complications in the form of localinfections, bacteremia, and sepsis. Patients who suffer from thesecomplications face a substantially longer hospital stay. Removal of theMCS 32 device requires “weaning” the patient from pump support by slowly(over hours) reducing pump speeds and closely monitoring hemodynamics toensure that the heart 12 does not decompensate from lack of support. Thedecision to begin weaning the patient from the MCS 32 device reliesheavily on a subjective clinical assessment of the patient (instead ofan objective measurement of CPO, impossible without native CO). In someresearch institutions, CPO is estimated directly using pressuremeasurement, and transthoracic echocardiography to look at native CO(Ejection Fraction) during the hours long weaning process, but thisrequires the resources of a surgeon, a heart failure cardiologist, andan expert cardiologist experienced in echocardiographic parameters. Thisis a huge amount of resources if the weaning process takes a long time,or if recovery is not actually complete. One potential clinical use ofmeasuring Native CO using the Abiomed Impella device, is that it canprovide a real-time assessment of patient condition.

Clinical Use 2: Native CO for automatic weaning control input—Weaning apatient from an implanted MCS 32 device need not be a manual process.Native CO can be measured in real time using the invention describedalongside pump flow to provide insight into the slowly improvingcondition of a patient supported by MCS 32. Weaning would happen in thisway optimally, allowing for total CO to remain constant, while thepatient's heart 12 recovers on its own. A number of well known-controlalgorithms could be applied automatically to the pump flow control basedon the patient's measured, Native CO.

Clinical Use 3: If pressure is available, then the device is capable ofmeasuring native pressure volume loops. Pressure-volume loops providevaluable information about the hemodynamic status of the heart.

Embodiment 1: Used in Preliminary Studies

Electrodes 18 used: The electrodes 18 used in this embodiment werepurchased in two sizes, to accommodate the two different catheter bodysizes, from Johnson Matthey, UK. Two were used as electrodes 1 and 2(with a smaller diameter), and two were used as electrodes 3 and 4 (witha larger diameter). The final embodiment is the bottom electrodeconfiguration.

Electrode Spacing: Electrode spacing was chosen to maximize the span ofthe chamber of interest (the left ventricle) while ensuring that allfour electrodes 18 would stay below the valve (and therefore in thechamber of interest). This was done by making sure that the mostproximal electrode (electrode 4) is close to the radiopaque marker 30that surgeons use to place the pump inlet inside the LV, and the pumpoutlet outside the LV (in the aorta). By co-locating electrodes 3 and 4with the radiopaque marker 30, it is ensured that the electrodes 18 willbe within the LV if the pump is functioning correctly.

Electrode Wiring 24: Wiring 24 considerations for the preliminarystudies embodiment required us to keep the wiring 24 on the outside ofthe lumen for the distal electrodes 1 and 2. This was done to keep thelumen free and clear for use by a guidewire to implant the catheter. Thewiring 24 on the proximal electrodes 3 and 4 is run internal to thecatheter body to avoid changing their large surface area, and the wiring24 on the distal electrodes 1 and 2 was run outside of the catheter withminimal contact area (see FIGS. 3, 6, 9 and 10). The key considerationfor the diameter (or gauge) of the wire is that they should be as thinas possible. If wiring 24 is run internal to the lumen between the pumpinlet and outlet, the requirement to be thin is to reduce impediment ofblood flow or required guidewires for implant. If wiring 24 is runexternal to the pump, then the wiring 24 is required to be thin toreduce the size of the sheath necessary to implant (allowing for aneasier surgical technique).

In another embodiment, as shown in FIG. 11, the wiring 24 runs internalto the plastic body (but not in the lumen space). This will be similarin function to the way that nitinol is embedded in the catheter bodyitself for the purpose of making the Impella catheter stiffer. Atradeoff in volume sensitivity to allow the electrodes 18 to be placedcloser together will also be exploited.

Motor 26 noise quantification: Fourier frequency analysis of the motor26 noise signal during pump operation was used to determine the optimalfrequencies to use for the cardiac volume measurements. It wasdetermined that the pump noise was at 50 khz, and at 100 kHz at all 9pump speeds for an Abiomed Impella device (Danvers, Mass.), and did notoverlap with the measurement. It was found from FIG. 7 that anyfrequency at least 5 kHz away from DC, and not an integer multiple of 50kHz could be utilized for the admittance-based volume measurement. 20kHz was chosen. In a fictitious example where 30 kHz was where motor 26noise was detected, any frequency from 10 kHz to 25 kHz, or from 35 kHzto 100 kHz, for example, can be chosen. FIG. 7 is a Fourier analysis ofmotor 26 electromagnetic noise vs. admittance signal for all 9 motor 26speeds (P1-P9). Note that the signal is 40 dB higher than thesurrounding noise floor (60 dB), meeting our success. Motor 26 noiseQuantification. Signal is at 20 kHz, note peaks at 50 and 100 kHzrepresenting motor 26 noise. Each color is a different pump speed.

WORKS CITED, ALL OF WHICH ARE INCORPORATED BY REFERENCE, HEREIN

-   [1] D. E. Haines et al., “Validation of a defibrillation lead    ventricular volume measurement compared to three-dimensional    echocardiography,” Heart Rhythm, June 2017.-   [2] L. M. Holt, M. L. Oglesby. A. P. Wang, J. W. Valvano, and M. D.    Feldman, “A Real-Time Hemodynamic ICD Measurement: Evaluation in    Chronically Implanted Canines With Pacing-Induced Dilated    Cardiomyopathy,” J Am Coll Cardiol EP, vol. 5, no. 6, pp. 742-743,    June 2019 PMID: 31221363.-   [3] E. R. Larson, M. D. Feldman, J. W. Valvano, and J. A. Pearce,    “Analysis of the Spatial Sensitivity of Conductance/Admittance    Catheter Ventricular Volume Estimation,” IEEE Trans. Biomed. Eng.,    vol. 60, no. 8, pp. 2316-2324, August 2013.-   [4] K. Raghavan et al., “Electrical Conductivity and Permittivity of    Murine Myocardium,” IEEE Trans. Biomed. Eng., vol. 56, no. 8, pp.    2044-2053, August 2009.

U.S. PATENTS AND APPLICATION

-   U.S. Pat. No. 10,420,952-   U.S. Pat. No. 10,376,177-   U.S. Pat. No. 10,076,669-   U.S. Pat. No. 9,820,673-   U.S. Pat. No. 9,295,404-   U.S. Pat. No. 7,925,335-   U.S. Pat. No. 6,494,832-   20110152661

Although the invention has been described in detail in the foregoingembodiments for the purpose of illustration, it is to be understood thatsuch detail is solely for that purpose and that variations can be madetherein by those skilled in the art without departing from the spiritand scope of the invention except as it may be described by thefollowing claims.

1: An apparatus for a heart of a patient comprising: a cardiac assistdevice adapted to be implanted into the patient to assist the heart withpumping blood, the cardiac assist device has a shaft that is adapted tobe positioned in the heart, the cardiac assist device includes a motorand an impeller disposed in the shaft which is driven by the motor toassist the heart with pumping blood, the pump draws blood from a leftventricle of the heart through an inlet port of the shaft and expelsblood into an ascending aorta of the heart through an outlet port of theshaft, thereby reducing some mechanical load on the heart and promotingrecovery; and a sensor adapted to be implanted into the patient, thesensor in communication with the cardiac assist device and the heartwhich measures native volume of the heart, the sensor includeselectrodes directly attached to the cardiac assist device that producesignals which are used to measure the native volume of the heart, andthe electrodes are in direct contact with the shaft that is positionedin the heart, the sensor includes a computer for data acquisition andanalysis of the signals, the computer in communication with theelectrodes, the computer produces the signals which include electricalcurrents to the electrodes and measures corresponding voltages to makeadmittance-based measurements and analyze the admittance-basedmeasurements to make real-time admittance-based volume measurements ofthe heart. 2-4. (canceled) 5: The apparatus of claim 16 wherein thecomputer provides electrical currents to the electrodes and measurescorresponding voltages to make admittance-based measurements and analyzethe admittance-based measurements to make real-time admittance-basedvolume measurements of the heart, the real-time admittance-based volumemeasurements of the heart by the computer include real-time measurementof both blood and muscle contributions. 6: The apparatus of claim 5wherein the sensor includes wiring that is in direct contact with theelectrodes and which extends to the computer over which the electricalcurrents and corresponding voltages pass.
 7. (canceled) 8: The apparatusof claim 6 wherein the cardiac assist device has a marker to guideproper placement of the cardiac assist device in the heart. 9: Theapparatus of claim 8 wherein the cardiac assist device is a temporarymechanical circulatory support (MCS) device which is a catheter-mountedblood pump that draws blood from a left ventricle of the heart throughan inlet port of the MCS and expels blood into an ascending aorta of theheart, thereby reducing some of the mechanical load on the heart andpromoting recovery. 10: An apparatus for a heart of a patientcomprising: a cardiac assist device adapted to be implanted into thepatient to assist the heart with pumping blood, the cardiac assistdevice has a shaft that is adapted to be positioned in the heart, thecardiac assist device includes a motor and an impeller disposed in theshaft which is driven by the motor to assist the heart with pumpingblood, the pump draws blood from a left ventricle of the heart throughan inlet port of the shaft and expels blood into an ascending aorta ofthe heart through an outlet port of the shaft, thereby reducing somemechanical load on the heart and promoting recovery; a sensor adapted tobe implanted into the patient, the sensor in direct contact with theshaft of the cardiac assist device and in communication with the heartwhich monitors the heart based on admittance while the cardiac assistdevice is in operation; and a pressure sensor adapted to be implantedinto the patient, the pressure sensor in contact with the shaft of thecardiac assist device and in communication with the heart and the heartwhich monitors left ventricular pressure while the cardiac assist deviceis in operation and with the computer plots native pressure volume loopswhile the cardiac assist device is in operation, the sensor includeselectrodes directly attached to the cardiac assist device that producesignals which are used to measure the native volume of the heart and theelectrodes are in direct contact with the shaft that is positioned inthe hear the sensor includes a computer for data acquisition andanalysis of the signals, the computer in communication with theelectrodes, the computer produces the signals which include electricalcurrents to the electrodes and measures corresponding voltages to makeadmittance-based measurements and analyze the admittance-basedmeasurements to make real-time admittance-based volume measurements ofthe heart. 11: An apparatus for a heart of a patient comprising: acardiac assist device adapted to be implanted into the patient to assistthe heart with pumping blood, the cardiac assist device has a shaft thatis adapted to be positioned in the heart, the cardiac assist deviceincludes a motor and an impeller disposed in the shaft which is drivenby the motor to assist the heart with pumping blood, the pump drawsblood from a left ventricle of the heart through an inlet port of theshaft and expels blood into an ascending aorta of the heart through anoutlet port of the shaft, thereby reducing some mechanical load on theheart and promoting recovery; a sensor adapted to be implanted into thepatient, the sensor in direct contact with the shaft of the cardiacassist device and in communication with the heart which monitors theheart based on impedance while the cardiac assist device is inoperation; and a pressure sensor adapted to be implanted into thepatient, the pressure sensor in contact with the shaft of the cardiacassist device and in communication with the heart and which monitorsleft ventricular pressure while the cardiac assist device is inoperation and with the computer plots native pressure volume loops whilethe cardiac assist device is in operation. 12: A method for treating aheart of a patient comprising the steps of: pumping blood of the patientwith a cardiac assist device implanted into the patient, the cardiacassist device has a shaft that is adapted to be positioned in the heart,the cardiac assist device includes a motor and an impeller disposed inthe shaft which is driven by the motor to assist the heart with pumpingblood, the pump draws blood from a left ventricle of the heart throughan inlet port of the shaft and expels blood into an ascending aorta ofthe heart through an outlet port of the shaft, thereby reducing somemechanical load on the heart and promoting recovery; and measuringnative volume of the heart with a sensor implanted into the patient, thesensor in communication with the cardiac assist device and the heart,the sensor includes electrodes directly attached to the cardiac assistdevice that produce signals which are used to measure the native volumeof the heart, and the electrodes are in direct contact with the shaftthat is positioned in the heart, the sensor includes electrodes directlyattached to the cardiac assist device that produce signals which areused to measure the native volume of the heart, and the electrodes arein direct contact with the shaft that is positioned in the heart, thesensor includes a computer for data acquisition and analysis of thesignals, the computer in communication with the electrodes, the computerproduces the signals which include electrical currents to the electrodesand measures corresponding voltages to make admittance-basedmeasurements and analyze the admittance-based measurements to makereal-time admittance-based volume measurements of the heart. 13: Theapparatus of claim 1 wherein the currents have a frequency and thesensor dynamically shifting the frequency of the currents to avoid noisein the patient to the sensor. 14: The apparatus of claim 13 wherein thecomputer of the sensor dynamically shifting the frequency of thecurrents to avoid noise from the cardiac assist device. 15: Theapparatus of claim 14 wherein the frequency of the currents includecurrents at a desired frequency and the computer includes a signalgenerator which generates the currents at the desired frequency andcauses changes in the frequency of the currents when dynamicallyshifting the currents. 16: The apparatus of claim 15 a pressure sensoradapted to be implanted into the patient, the pressure sensor in contactwith the shaft of the cardiac assist device and in communication withthe heart and the heart which monitors left ventricular pressure whilethe cardiac assist device is in operation and with the computer plotsnative pressure volume loops while the cardiac assist device is inoperation.